Biocompatible and biodegradable scaffolds are used in tissue reconstruction and repair. The scaffold, for example, can serve as both a physical support and adhesive substrate for cells during in vitro culturing and subsequent implantation in vivo. Scaffolds are utilized to deliver cells to desired sites in the body, to define a potential space for engineered tissue, and to guide the process of tissue development. Cell transplantation on scaffolds has been explored for the regeneration of skin, heart, nerve, liver, pancreas, cartilage, and bone tissue using various biological and synthetic materials. Scaffolds have also been implanted directly into patients without prior culturing of cells in vitro. In this case, the initially cell-free scaffold needs to be designed in a manner that cells from the surrounding living tissue can attach to the scaffold, migrate into it and form functional tissue.
Synthetic materials, such as polyester fiber (e.g., DACRON™) or polytetrafluororethylene (PTFE) (e.g., TEFLON™), have been extensively used as implants to replace diseased or damaged body parts. However, these materials have enjoyed limited success because of their poor biocompatibility. Synthetic materials frequently cause persistent inflammatory reactions and are further limited by their lack of biodegradation, which results in the failure of the body to integrate and/or remodel these materials into the surrounding tissue. Other synthetic materials that are more biodegradable than polyester fibers or PTFE have been used to fabricate tissue-engineering scaffolds such as polyglycolic acid (PGA), polylactic acid (PLA). PGA, PLA and their copolymers are the most commonly used synthetic polymers in tissue engineering. However, in order for these structures to promote functional tissue growth/reconstruction, a complex micro-architecture must be produced, and this requires difficult processing methods.
Non-synthetic materials, such as animal materials, have also been used to produce tissue-regeneration/repair scaffolds. Efforts to use animal materials have been unsatisfactory when these materials are cross-linked by formaldehyde or glutaraldehyde, for example. This form of generalized aldehydic crosslinking tends to render biomaterials sufficiently unrecognizable to tissue cells such that normal remodeling and integration are not promoted. Similarly, other types of chemical processing of animal or human biomaterials, such as extraction with detergents, hypertonic buffers or hypotonic buffers can alter them to the degree that leave these biomaterials toxic to tissue cells, ineffective in promoting angiogenesis, and ineffective in stimulating repair and remodeling processes needed for the conversion of an implant into a functional substitute for the tissue or organ being replaced.
Another approach uses extracellular matrix (ECM) components in processed or natural forms to regenerate tissue in vitro and in vivo. The interaction of cells with ECM in in vivo and in vitro environments is important in the organization, function and growth of all tissues and organs. Biochemical and biophysical signals between the cell and the ECM regulate fundamental cellular activities including adhesion, migration, proliferation, differential gene expression, and programmed cell death. Processed forms of ECM have been used as tissue regeneration aids. For example, as described in U.S. Pat. No. 5,275,826, fluidized intestinal submucosa can be injected into host tissues in need of repair or used in combination with other graft materials. Basement-membrane-derived ECM compositions (e.g., MATRIGEL™, BECTON-DICKINSON) in the form of polymerizable extracts have also been used to regenerate or repair tissue. These polymerizable extracts can be formed or molded into a three-dimensional gel structure attempting to resemble lamellar structures (U.S. Pat. No. 4,829,000). Collagen-based gels have also been combined with specialized cells. This process depends upon interactions between the cells and collagen filaments in the gel so that the cells condense and organize. While tissue-like constructs have been fabricated and been shown to have some resemblance to their natural counterparts, these constructs do not readily develop the matrix complexity characteristic of the actual tissues that they are meant to imitate (U.S. Pat. Nos. 4,485,096 and 4,485,097).
Natural forms of ECM (e.g., isolated sheets or layers) can be obtained from tissue submucosa of warm-blooded vertebrates. For example, the tunica submucosa of the intestine is often used as tissue graft material; see U.S. Pat. Nos. 4,902,508 and 5,281,422. Both stomach (U.S. Pat. No. 6,099,567) and urinary bladder (U.S. Pat. Nos. 5,554,389 and 6,171,344) submucosa have also been described as natural sources of ECM. Large sheets of submucosal tissue can be prepared from smaller segments of submucosal tissue through conventional techniques such as weaving, knitting or the use of adhesives (U.S. Pat. No. 5,997,575). Purified submucosa can also be shaped into other forms with other graft materials, for example, a tubular graft composed of sheets of submucosa wrapped around a tube (U.S. Pat. No. 6,358,284). However, a limitation of these natural forms is that it is difficult to fashion three-dimensional tissue structures that mimic the endogenous tissue architecture where the graft or scaffold is to be used, for example, to form the complex shape of a valve.
Cardiac tissue repair presents challenges to tissue-engineering methods. Cardiac tissues are difficult to repair as the tissue has a limited ability to regenerate and because the tissue has a complex architecture that is difficult to mimic. For example, cardiomyocytes have a natural complex architecture with particular electromechanical properties. Endogenous cardiomyocytes are organized into parallel cardiac muscle fibers with intracellular contractile myofibrils oriented parallel to the long axis of each cell. Junctional complexes between abutting cells are concentrated at the ends of each cardiomyocyte. Such architecture provides the electromechanical coupling of cardiomyocytes needed to stimulate the transmission of directed contraction over long distances. However, cultured cardiomyocytes typically spread on a flat substrate to form an unorganized monolayer with disorganized myofibrils and junctions. In vitro methods have been created that align cardiomyoctes into functional organizations by growing cultured cardiomyocytes on flat micro-patterned substrates (McDevitt, T. C. et al., “In vitro generation of differentiated cardiac myofibers on micropatterned laminin surfaces,” J. Biomed. Mater. Res., (2002), 60:472-479). However, such methods provide two-dimensional tissue layers that are difficult to transform into a three-dimensional tissue repair scaffold.
Thus, biocompatible, biodegradable, three-dimensional scaffolds that are adaptable to different in vivo tissue architectures and that promote the natural, functional coupling between cells present in the scaffold and cells in the surrounding tissue are desired. Further, methods of producing such improved scaffolds that are efficient and simple are also desired.